This invention relates to a nuclear magnetic resonance imaging apparatus.
FIG. 1 is a cut-away perspective view of a conventional nuclear magnetic resonance imaging apparatus of the type to which the present invention pertains. As shown in this figure, superconducting coils 1 which generate a high-uniformity static magnetic field are surrounded by cylindrical thermal shield 2 which shields the superconducting coils 1 from heat transfer. The thermal shield 2 is made of a metal of good electrical conductivity such as copper or aluminum. The superconducting coils 1 and the thermal shield 2 together constitute a superconducting magnet and are disposed inside a cylindrical housing 3. The inner periphery of the housing 3 surrounds a cylindrical frame 6 on which a plurality of magnetic field coils 4a, 4b, 5a, and 5b are wound. Magnetic field coil 4a generates a gradient field in the direction of the X axis in FIG. 1, and magnetic field coil 4b generates a gradient field in the direction of the Y axis. Magnetic field coils 5a and 5b are in the form of coaxial rings and will hereinafter be referred to as circular coils. The circular coils 5a and 5b generate a gradient field in the direction of the Z axis when current is passed through the coils in opposite directions. A human patient 7 who is used as a subject for imaging lies inside the apparatus upon a movable bed 8 by means of which his position can be adjusted.
In a nuclear magnetic resonance imaging apparatus of this type, imaging control of the nuclear magnetic resonance which is induced in the patient 7 is performed using the high-uniformity static magnetic field which is generated by the superconducting coils 1 and the gradient field which are generated in the X, Y, and Z directions, respectively, by the magnetic field coils 4a, 4b, and 5a and 5b. It is important that the magnetic flux density of the magnetic field in the direction of the Z axis should vary as linearly as possible along the Z axis. For this reason, the circular coils 5a and 5b are positioned in accordance with the following principles, which will be explained while referring to FIG. 2, which is a schematic perspective view of the circular coils 5a and 5b of FIG. 1. The origin O of the coordinate axes in the figure lie midway between the two circular coils 5a and 5b, and the Z axis coincides with a line connecting their centers. If the radius of both circular coils 5a and 5b is a.sub.1 and the Z coordinates of the centers of the coils are +Z.sub.1 and -Z.sub.1, respectively, the magnetic flux density B(Z) of the magnetic field at an arbitrary point Z along the Z axis when a current I is passed through the coils 5a and 5b in opposite directions is given by the following equation: ##EQU1##
Where, .mu..sub.0 is the permeability of vacuum and .epsilon..sub.1 (.beta.), .epsilon..sub.3 (.beta.), and .epsilon..sub.5 (.beta.) are functions of the ratio .beta.=Z/a.sub.1, i.e., the ratio between the distance Z.sub.1 of the centers of the circular coils from the origin O and the coil radius a.sub.1. In order for the magnetic flux density B(Z) which is given by Equation (1) to be linearly proportional to Z, it is necessary only to eliminate the term proportional to Z.sup.3 and all higher order terms. The value of the function .epsilon..sub.3 (.beta.) is given by the following equation. ##EQU2##
Accordingly, if .beta. is chosen to be .+-.3/2, then .epsilon..sub.3 (.beta.)=0, and the Z cubed term in Equation (1) is eliminated. Furthermore, when Z/a.sub.1 &lt;1, the term proportional to (Z/a.sub.1).sup.5 is small compared to the term proportional to Z and are negligible. Therefore, B is set equal to .+-..sqroot.3/2. In other words, Z.sub.1 =.+-..sqroot.3a.sub.1 /2, and in the coil system of FIG. 2, in the region of interest near the origin O, a magnetic field can be generated whose magnetic flux density is linearly proportional to the Z coordinate.
However, in a conventional nuclear magnetic resonance imaging apparatus having the above-described geometry, a magnetic flux density of high linearity can be obtained only under ideal conditions in which there are no fluctuations due to surrounding magnetic fields. Under actual conditions, as the thermal shield 2 is a metal of good electrical conductivity, when pulse currents which are necessary for nuclear magnetic resonance imaging flow through the circular coils 5a and 5b, eddy currents are induced in the thermal shield 2 due to the magnetic field which is generated by the circular coils 5a and 5b. The resulting magnetic field along the Z axis is therefore a composite magnetic field, i.e., a combination of the magnetic fields due to the eddy currents and the magnetic field generated by the circular coils 5a and 5b. Because of the component due to the magnetic fields generated by the eddy currents, the composite magnetic field does not vary linearly along the Z axis.